Implantable stimulation devices are devices that generate and deliver electrical stimuli to body nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder sublaxation, etc. The present invention may find applicability in all such applications, although the description that follows will generally focus on the use of the invention within a Spinal Cord Stimulation (SCS) system, such as that disclosed in U.S. Pat. No. 6,516,227 (“the '227 patent”).
Spinal cord stimulation is a well-accepted clinical method for reducing pain in certain populations of patients. As shown in FIGS. 1A and 1B, a SCS system typically includes an Implantable Pulse Generator (IPG) 100, which includes a biocompatible case 30. The case 30 usually holds the circuitry and power source or battery necessary for the IPG to function. The IPG 100 is coupled to electrodes 106 via one or more electrode leads (two such leads 102 and 104 are shown), such that the electrodes 106 form an electrode array 110. The electrodes 106 are carried on a flexible body 108, which also houses the individual signal wires 112, 114, coupled to each electrode. The signal wires 112 and 114 are connected to the IPG 100 by way of an interface 115, which may be any suitable device that allows the leads 102 and 104 (or a lead extension, not shown) to be removably connected to the IPG 100. Interface 115 may comprise, for example, an electro-mechanical connector arrangement including lead connectors 38a and 38b configured to mate with corresponding connectors 119a and 119b on the leads 102 and 104. In the IPG 100 illustrated in FIG. 1A, there are eight electrodes on lead 102, labeled E1-E8, and eight electrodes on lead 104, labeled E9-E16, although the number of leads and electrodes is application specific and therefore can vary. The electrode array 110 is typically implanted along the dura of the spinal cord, and the IPG 100 generates electrical pulses that are delivered through the electrodes 106 to the nerve fibers within the spinal column. The IPG 100 itself is then typically implanted somewhat distantly in the buttocks of the patient.
As shown in FIG. 2, an IPG 100 typically includes an electronic substrate assembly 14 including a printed circuit board (PCB) 16, along with various electronic components 20, such as microprocessors, integrated circuits, and capacitors, mounted to the PCB 16. Ultimately, the electronic circuitry performs a therapeutic function, such as neurostimulation. A feedthrough assembly 24 routes the various electrode signals from the electronic substrate assembly 14 to the lead connectors 38a, 38b, which are in turn coupled to the leads 102 and 104 (see FIGS. 1A and 1B). The IPG 100 further comprises a header connector 36, which among other things houses the lead connectors 38a, 38b. The IPG 100 can further include a telemetry antenna or coil (not shown), which can be mounted within the header connector 36, for receipt and transmission of data to an external device such as a hand-held or clinician programmer (not shown). As noted earlier, the IPG 100 usually also includes a power source, typically a rechargeable battery 26.
Also shown in FIG. 2 is an external charger 12 that is used to provide power to the IPG 100, which is explained in further detail below. The external charger 12 itself needs power to operate, and therefore may include its own battery 70, which may also be a battery that is rechargeable using a plug-in-the-wall holster (“cradle”) or power cord connection much like a cellular telephone. Alternatively, the external charger 12 may lack a battery and instead draw its power directly from being plugged into a wall outlet (not shown). In any event, a primary function of the charger 12, as discussed further below, is to energize a charging coil 17. The external charger 12 can contain one or more circuit boards 72, 74, which contain the circuitry 76 needed to implement such functionality. In a preferred embodiment, and as shown in FIG. 2, most of the circuitry 76 can be located on an orthogonal circuit board 74, which reduces interference and heating that might be produced by the charging coil 17, as is further explained in U.S. Patent Application Publication No. US 2008/0027500.
Further details concerning the structure and function of typical IPGs and IPG systems are disclosed in U.S. Pat. No. 7,444,181.
If the battery 26 in the IPG 100 is rechargeable, it will be necessary to charge the battery 26 periodically using the external charger 12, i.e., a charger that is external to the patient in whom the IPG 100 is implanted. Because the IPG 100 may already be implanted in a patient, wireless recharging is greatly preferred to obviate the need to replace a power-depleted battery 26 via surgery.
To convey energy wirelessly between the external charger 12 and the IPG 100, and as shown in FIG. 2, the charger 12 typically includes an energized alternating current (AC) coil 17 that supplies energy 29 to a similar charging coil 18 located in or on the IPG 100 via inductive coupling. In this regard, the coil 17 within the external charger 12 is wrapped in a plane 50, which lies substantially parallel to the plane 52 of the coil 18 within the IPG 100, as shown schematically in FIG. 3. Such a means of inductive energy transfer can occur transcutaneously, i.e., through the patient's tissue 25. The energy 29 received by the IPG 100's coil 18 can then be rectified and stored in a rechargeable battery 26 within the IPG 100, which in turn powers the electronic circuitry that runs the IPG 100. Alternatively, the energy 29 received can be used to directly power the IPG 100's electronic circuitry, which may lack a battery altogether.
Conventional external chargers 12 typically employ relatively simple user interfaces 94, which simplicity is warranted either because of the relative simplicity of the charging function, or because the external charger 12 may not be visible to the patient while in use, which limits the utility of more complex visual user interfaces. For example, in an SCS application in which the IPG 100 is typically implanted in the buttocks, the external charger 12 is generally behind the patient while charging to align the external charger 12 with the IPG 100. Additionally, the external charger 12 may be covered by clothing while in use, again reducing the utility of a visual user interface. The user interface 94 of the conventional external charger 12 of FIG. 2 therefore typically merely comprises an on/off switch that activates the charger 12, an LED to indicate the status of the on/off switch, and a speaker for emitting a “beep” at various times, such as when the charger is not properly aligned with the IPG 100 or when charging has completed.
Inductive charging between the two coils 17, 18 can produce significant heating in the external charger 12. Such external charger heating could, if unchecked, possibly discomfort or injure the patient. This possibility of injury is heightened because the external charger 12 is often held against the patient's tissue 25 during charging. For example, in an SCS system, the external charger 12 is generally held in place against the buttocks of the patient by a “fanny pack.”
Accordingly, prior art external chargers have incorporated temperature monitoring and control circuitry to detect external charger temperatures, and to control charging accordingly. For example, and as shown in FIG. 2, a prior art external charger 12 can include one or more temperature sensors 92, which for example can comprise thermistors or thermocouples affixed by heat conducting epoxy to the housing of the external charger 12. A hole 90 in the circuit board 72 can assist in connecting the temperature sensor 92 to the temperature sensing circuitry (not shown) resident on either of circuit boards 72 or 74. The temperature monitoring and control circuitry generally senses the temperature, T(EC), of the external charger, and in particular sets a maximum temperature, Tmax(EC) for the external charger. The maximum temperature Tmax(EC) may be set to 41° C. (˜106° F.) for example, which temperature is conservatively picked by the manufacturer of the external charger 12 as a temperature that should not discomfort or injure a normal healthy adult.
The temperature monitoring and control circuitry in the external charger 12 can operate as illustrated in FIG. 4, which shows the temperature of the external charger, T(EC), during a typical charging session. Initially, the charging circuitry in the external charger 12 is enabled, i.e., an AC current flows through coil 17 in the external charger 12 as previously discussed. As this occurs, T(EC) increases. Eventually, T(EC) equals Tmax(EC). At this point, the temperature sensing circuitry would inform the microcontroller in the external charger 12 to suspend charging, i.e., to cease current flow through coil 17. Once the current ceases, the T(EC) will start to fall. At some point—for example after some time duration or when a minimum T(EC) (Tmin(EC)) is reached as illustrated—charging can be enabled until once again T(EC) reaches Tmax(EC), etc. The result is that charging is duty cycled between enabled and disabled states.
Although the charging scheme illustrated in FIG. 4 ensures that the external charger 12 never exceeds a predefined maximum safe temperature, Tmax(EC), the inventors consider such scheme non-optimal, because it fails to allow for differences between patients, and does not provide any way to control external charger 12 heating characteristics. For example, if a patient is not particularly heat sensitive, that patient may be able to tolerate a higher Tmax(EC), such as 42° C. for example. However, if Tmax(EC) is constrained to 41° C. by the manufacture of the external charger 12, charging will not take place as aggressively as that patient could tolerate: the current in the charging coil 17 would be limited, or charging would be suspended for greater amounts of time. In either case, the result is that charging will be performed too slowly for that patient. This is inconvenient, as patients would generally like charging to occur as quickly as possible. On the other hand, if a patient is unusually heat sensitive for some reason, perhaps because of a medical condition, that patient might be more comfortable with a lower Tmax(EC), say 40° C. for example. In this case, if Tmax(EC) were constrained to 41° C. by the manufacture, that patient would perceive charging as uncomfortably warm.